1. Field of the Invention
This invention relates to magnetic resonance imaging (MR imaging), and more particularly to computing quantitative images as well as synthetic images from scan data, with the purpose of allowing the user to perform virtual MRI scanning retrospectively and not requiring the presence of the patient.
2. Background Information
In magnetic resonance image scanning (MRI scanning), images of a subject, usually a patient""s body, are produced through the interaction of a magnetic field applied to the patient""s body and the magnetic moment of protons. Each proton behaves as small bar magnet, and the strength of the bar magnet is referred to as the xe2x80x9cmagnetic momentxe2x80x9d of the proton. All protons have the same value of magnetic moment, just as each proton has the same value of electric charge. The protons are the nuclei of hydrogen atoms. The hydrogen is chemically bonded in compounds of the patient""s tissue.
It is standard engineering practice in MRI imaging to apply a strong magnetic field substantially parallel to the spinal column of the patient. This magnetic field is referred to as the xe2x80x9clongitudinal magnetic fieldxe2x80x9d and is represented in symbols as B0. Upon the application of the longitudinal magnetic field, protons in the patient""s tissue align with the magnetic field to produce a magnetization (longitudinal magnetization) of the patient""s tissue. The longitudinal magnetization is a vector quantity that points along the applied longitudinal magnetic field. The magnetization of the patient""s tissue may be represented as formed by many protons aligned with the longitudinal magnetic field.
The patient""s magnetization is used to produce images by observing its response to a magnetic field applied by radio frequency pulses, where the radio frequency magnetic field is applied perpendicular to the longitudinal magnetic field. The frequency of the radio frequency magnetic field is chosen, along with the time duration of the application of the radio frequency magnetic field, to cause the protons to rotate, more precisely to precess, by a desired angle. The angle by which the protons rotate, or precess, is referred to as the xe2x80x9cflip anglexe2x80x9d. Ordinarily, radio frequency magnetic fields are applied to rotate the protons through an angle of 180 degrees so that their magnetization points in the reverse direction of the applied longitudinal magnetic field (that is into the anti-parallel direction), or to rotate the protons through an angle of 90 degrees so that their magnetization points into a plane perpendicular to the direction of the applied longitudinal magnetic field, that is into the xe2x80x9ctransverse planexe2x80x9d. Other values of the flip angle may also be employed in MRI imaging.
An image reproducing the proton density in the patient""s body may be obtained by applying a radio frequency field for a time sufficient to rotate the proton magnetization through to a 90-degree angle, and such an application of a radio frequency magnetic field is referred to as applying a xe2x80x9c90 degreexe2x80x9d RF pulse. Upon application of a 90 degree RF pulse, the protons are rotated from a direction substantially parallel to the longitudinal magnetic field into a direction substantially perpendicular to the longitudinal magnetic field. The proton magnetic moments, during this rotation, remain substantially aligned with each other, so the patient magnetization becomes a vector in the plane perpendicular to the longitudinal magnetic field. The plane perpendicular to the longitudinal magnetic field is referred to as the xe2x80x9ctransverse planexe2x80x9d.
The patient""s magnetization in the transverse plane is substantially equal in magnitude to the value that it had before application of the 90 degree RF pulse, however the patient""s magnetization points in a direction in the transverse plane. For example, the transverse plane can be described with an X-axis and a Y-axis, and the X-axis may be chosen so that it is aligned with the magnetization in the transverse plane at the end of the 90-degree RF pulse. The magnetization in the transverse plane rotates in the transverse plane, and as a consequence of this rotation generates a radio frequency signal originating from the patient""s tissue. This radio frequency signal is detected by a radio receiver, and is analyzed to produce an image.
A particular transverse plane is chosen for readout by applying a longitudinal magnetic field gradient, and choosing the frequency of the RF pulse to resonate with the protons in the chosen transverse plane. Ordinarily, changing the frequency of the RF pulse while longitudinal magnetic field gradient is held constant shifts the position of the desired transverse plane. The radio receiver receives the RF signal generated by the rotating magnetization, and the RF signal received by the radio receiver is spread over a frequency band, and with different phases, by the application of two transverse magnetic field gradients. For example, a transverse magnetic field gradient is applied, and a read out of emissions from the patient""s tissue is obtained from the RF receiver. Again, a different transverse magnetic field gradient is applied, and a second readout is obtained from the RF receiver. A sequence of readouts is obtained, for different radio frequency values and for different phases, by applying different magnetic field gradients. A Fourier transform of the frequency and phase information received by the RF receiver is then computed. The output of the Fourier transform calculation produces the image of the patient""s tissue. The image is presented as a two dimensional matrix of pixels.
After a first image is obtained, a waiting period is introduced. At the end of the waiting period a second image of the patient""s tissue in the same transverse plane is obtained. The intensity of the radio frequency signal generated by the patient""s tissue is reduced in the second image in comparison with the strength produced in the first image by transverse relaxation phenomena. Transverse relaxation phenomena are modeled by a transverse relaxation time, referred to as xe2x80x9cT2xe2x80x9d. This reduction in the signal in the second image is used to compute the transverse relaxation time T2. The values of T2 may be computed at each pixel of the image.
Transverse relaxation phenomena, which are measured by the measured value of T2, are predominately caused by different protons in the transverse plane being subject to slightly different magnetic fields. The different magnetic fields throughout the transverse plane have their origin in several phenomena: the first being the magnetic field gradient which is intentionally applied to the transverse plane in order to obtain space resolution in the transverse plane; a second being different chemical environments of the protons in molecules within different regions of the transverse plane; and a third being movement of the protons through the tissue of the patient, such as caused by blood flow, etc.
Additionally, the protons in the transverse plane relax back to being parallel to the longitudinal magnetic field with a relaxation time referred to as xe2x80x9cT1xe2x80x9d, where T1 is known as the xe2x80x9clongitudinal relaxation timexe2x80x9d.
A 180-degree RF pulse may be applied to the patient""s tissue. The result of the 180-degree RF pulse is that the proton magnetic moments are rotated 180 degrees. This rotation points the protons away from being parallel to the longitudinal magnetic field to being anti-parallel. The protons then relax toward the parallel orientation of the longitudinal magnetic field with the transverse relaxation time xe2x80x9cT1xe2x80x9d. After a waiting time from the 180-degree pulse, a 90-degree RF pulse is then applied to the patient""s tissue. The 90-degree RF pulse causes the net magnetization of the patient""s tissue to rotate 90 degrees into the transverse plane where the magnetization precesses, and so produces an output RF signal from the patient""s tissue. An image is again read out by using transverse magnetic field gradients and the RF receiver.
It is standard engineering practice to use a sequence of 180-degree RF pulses and 90-degree RF pulses in order to generate desired images by an MRI imaging apparatus.
In describing images of a patient""s body taken by MRI imaging, it is necessary to introduce some additional geometric terminology. The patient""s body is regarded as being made of slices, where the slices are in the transverse plane, that is in planes perpendicular to the longitudinal direction. The slices have thickness in the longitudinal dimension, for example a thickness of 0.5 millimeter, 1 millimeter, 2 millimeters, etc. (hereinafter millimeters will be abbreviated by xe2x80x9cmmxe2x80x9d). An xe2x80x9cxxe2x80x9d and xe2x80x9cyxe2x80x9d axis are defined as fixed in the patient""s body, and the xe2x80x9cxxe2x80x9d and xe2x80x9cyxe2x80x9d axes lie in a transverse plane, where a xe2x80x9czxe2x80x9d axis is introduced as lying along the longitudinal axis and parallel to the applied longitudinal magnetic field. Each of the slices is divided by lines parallel to the patient""s xe2x80x9cxxe2x80x9d and xe2x80x9cyxe2x80x9d axes to form small rectangular parallelepiped elements of patient""s tissue referred to as xe2x80x9cvoxelsxe2x80x9d. A voxel is a small volume element of patient""s tissue. The values of observed quantities in voxels are presented as a two-dimensional image, where each element in the two dimensional image is referred to as a xe2x80x9cpixelxe2x80x9d. Quantities, which may be observed in a voxel during MRI imaging, are: proton density; value of transverse relaxation time T2 in the voxel; value of longitudinal relaxation time T1 in the voxel, etc. The values of the quantities observed in a voxel are then presented as a two dimensional image of pixels, where the image is usually presented on a computer screen where it may be photographed, printed by a computer printer, etc.
For example, an analysis of two images following a 90-degree RF pulse yields a value of the transverse relaxation time T2. For example, after application of the 90 degree RF pulse, a waiting time, referred to as xe2x80x9cTE1xe2x80x9d is observed. After the waiting time TE1 is observed, a readout of an image is obtained. After a further waiting time referred to as xe2x80x9cTE2xe2x80x9d, a second image is obtained.
When a 90 degree RF pulse is applied to a patient, the first image that is taken after the TE1 waiting time is referred to as a proton density weighted image, a xe2x80x9cPD weightedxe2x80x9d image. The terminology xe2x80x9cPD weighted imagexe2x80x9d is used because this is the first image in a sequence, and the intensity of the image in each voxel of the image is regarded as being proportional to the proton density in the corresponding voxel of the patient""s transverse plane. The two-dimensional image of pixels generated from the intensity of the signal observed in the voxels is then presented on a computer screen, etc.
The second image produced after the TE2 waiting time is referred to as a xe2x80x9cT2 weightedxe2x80x9d image. The terminology xe2x80x9cT2 weightedxe2x80x9d image is used because the image is reduced in intensity at each voxel because of transverse relaxation processes, and so is dependent upon the value of T2 at each voxel of the transverse plane.
When a 180-degree RF pulse is applied to a patient, the longitudinal magnetization is rotated to an anti-parallel direction to the applied longitudinal magnetic field, and then the longitudinal magnetization begins relaxing back to the parallel direction with the relaxation time of T1. The value of T1 at each voxel may be determined by a sequence of images. Ordinarily the images to determine T1 are taken by: first, applying the 180 degree RF pulse to obtain anti-parallel orientation of the patient""s longitudinal magnetization; second, by applying a 90 degree RF pulse after a longitudinal waiting time of invT from the 180 degree RF pulse to rotate the longitudinal magnetization into the transverse plane; and third by taking one or more subsequent images in order to obtain the value of T2 at each voxel of the slice. Images taken after at least two different values of longitudinal waiting time invT may then be analyzed to obtain a value of T1 at each voxel. These two images will be of different intensity because of the recovery of longitudinal magnetization caused by the longitudinal relaxation processes modeled by the longitudinal relaxation time T1. And the value of T1 may be different in each voxel, giving rise to images having different contrast information at the different values of waiting time invT.
It is standard engineering practice to take the images to determine T1 by the additional steps of: first applying a 90 degree RF pulse after the waiting time of invT from the 180 degree RF pulse to rotate the longitudinal magnetization into the transverse plane, and then taking an image after a waiting time of TE1, and a second image after an additional waiting time of TE2, where these values of TE1 and TE2 are the same as the waiting times for taking an image after an isolated 90 degree RF pulse. The images taken after an isolated 90 degree RF pulse, and after a 180 degree pulse followed by a 90 degree RF pulse, may then be combined in an analysis to obtain values of T1 and T2 at each voxel of the patient""s body.
It is a standard clinical and engineering terminology to refer to: the first image after a 90 degree RF pulse as a xe2x80x9cPD weightedxe2x80x9d image (proton density image); the second image after a 90 degree RF pulse as a xe2x80x9cT2 weightedxe2x80x9d image; the first image after a 180 degree RF pulse and a waiting time of invT, then followed by a 90 degree pulse, as a xe2x80x9cT1 weightedxe2x80x9d image; and the second image after both the 180 degree RF pulse and the 90 degree RF pulse as a T1-T2 weighted image.
Clinical information is carried in the different images, the PD weighted image, the T1 weighted image, the T2 weighted image, the T1-T2 weighted image, and other images produced by a particular choice of RF pulse sequences. In order to produce images by different RF pulse sequences, it is necessary to place the patient in the MRI imaging scanner and to subject the patient to the desired sequence of RF pulses.
There is needed a way to improve the images obtained from the radio receiver data, so that clinical information may be obtained without regard to the sequence of RF pulses applied to the patient""s tissue.
The invention consists of three image-postprocessing phases for the purposes of generating high-quality quantitative MR images (proton density (PD), T1, and T2) as well as high-quality virtual MR images with continuously adjustable computer-synthesized contrast weightings, from source images acquired directly with an MRI scanner. Each of the image-postprocessing phases uses one or several new computer algorithms that improve image quality with respect to prior art, as described in detail in the following sections of this document.
Specifically, the Image-postprocessing Phases are:
1) Generation of high-quality synthetic proton density images by computing linear combinations of source images (LCSI algorithms) that were directly acquired with an MRI scanner. The objective of this phase is to generate a high-quality computer representation of the patient for the purpose of virtual MRI scanning, or equivalently for using this representation as a cybernetic virtual patient. The source images used as input to the LCSI algorithm(s), which are directly acquired with the MRI scanner, may include one, several, or all the images generated with the application of the physical pulse sequence used to scan the patient. Furthermore, for the purpose of generating a computer representation of the patient for virtual MRI scanning, the specific applied MRI pulse sequence(s) as well as the number or the individual contrast weightings of the source images per slice generated with it, are not unique. As discussed below in this patent application, the first phase of the invention (i.e. generating proton density images as linear combinations of source images) was applied to source images that were generated with a pulse sequence known as mixed turbo spin-echo (MIX-TSE) but many other are possible. The MIX-TSE pulse sequence used here as an example, produced four images per anatomic slice, specifically: a PD weighted image, a T1 weighted image, a T2 weighted image, and a combination T1-T2 weighted image. The values of the coefficients for the LCSI algorithms are also not uniquely defined. These values are chosen first and foremost so as to produce accurate representations of the proton density spatial distribution within the slice and second, so as to maximize the signal-to-noise-ratio (SNR) of the generated image. Two exemplary sets of LCSI algorithms are discussed in this document: 1) LCSI-basic: consist of a simple set of linear coefficients that uses solely the source image that is closest to the proton density image, and 2) LCSI-maximum-SNR: consists of a set of coefficients obtained by means of an optimization algorithm to maximize signal to noise ratio, SNR, which uses all source images (4 in the case of this example). Once the values of the LCSI coefficients are chosen, LCSI proton density images (LCSI_PD) are calculated at every location of patient that was physically scanned with the MIX-TSE pulse sequence.
2) In the second phase of the invention, the values of longitudinal relaxation time T1 and transverse relaxation time T2 are computed at each scanned voxel of the patient using also the source images and by means of what will be referred to hereafter as model-conforming quantitative algorithms. The defining feature of model-conforming Q-MRI algorithms is that these are based on a physics model of tissue magnetism that is conditioned in response to signal-to-noise-ratio criteria in the source images. Specifically, in the model conforming algorithms the pixel values of the input source images are first compared to the noise level in the image, then, if these pixel values exceed the noise level the physics model applies, otherwise a predefined value for the output (T1 and T2) results. The use of model conforming algorithms is shown to reduce the incidence of inaccurate pixel values in Q-MR images as well as in synthetic images produces by virtual MRI scanning.
3. In the third phase of the invention, the values of the LCSI-proton density, T1, and T2 obtained for each voxel are then used to compute a synthetic MR image with computer-generated contrast weighting. In this phase, synthetic images representing any imaginable sequence of pulses may be computed. Also, the voxels may be rearranged to give slices through the patient""s body at any desired orientation, for example slices parallel to the transverse plane, slices parallel to the longitudinal axis, and slices at any desired angle to the longitudinal axis and the transverse plane.
In summary, the qualityxe2x80x94high SNR, negligible incidence of pixel dropout artifacts, and continuously adjustable contrast weightingsxe2x80x94of the synthetic images produced with the algorithms of this invention, is consistently superior to that produced by prior art. These image quality improvements result primarily from using LCSI_PD images as the virtual patient, and secondarily from using model-conforming Q-MR images (e.g. T1 and T2) to produce contrast weighting in virtual MRI scanning.